Method of making a stent formed from crosslinked bioabsorbable polymer

ABSTRACT

A stent having a stent body made from a crosslinked bioabsorbable polymer is disclosed. A method of making the stent including exposing a tube formed from a bioabsorbable polymer to radiation to crosslink the bioabsorbable polymer and forming a stent body from the exposed tube is disclosed. The tube can include a crosslinking agent which induces crosslinking upon radiation exposure. Additionally or alternatively, the bioabsorbable polymer can be a copolymer that crosslinks upon exposure to radiation in the absence of a crosslinking agent.

This application is a continuation application of U.S. application Ser.No. 12/422,143 filed on Apr. 10, 2009, which is incorporated herein byreference.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to methods of manufacturing polymeric medicaldevices, in particular, stents.

2. Description of the State of the Art

This invention relates to radially expandable endoprostheses, that areadapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel. Astent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices that function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels.

“Stenosis” refers to a narrowing or constriction of a bodily passage ororifice. In such treatments, stents reinforce body vessels and preventrestenosis following angioplasty in the vascular system. “Restenosis”refers to the reoccurrence of stenosis in a blood vessel or heart valveafter it has been treated (as by balloon angioplasty, stenting, orvalvuloplasty) with apparent success. Stent are also used widely inendovascular applications, such as in the popliteal artery.

Stents are typically composed of scaffolding that includes a pattern ornetwork of interconnecting structural elements or struts, formed fromwires, tubes, or sheets of material rolled into a cylindrical shape.This scaffolding gets its name because it physically holds open and, ifdesired, expands the wall of the passageway. Typically, stents arecapable of being compressed or crimped onto a catheter so that they canbe delivered to and deployed at a treatment site.

Delivery includes inserting the stent through small lumens using acatheter and transporting it to the treatment site. Deployment includesexpanding the stent to a larger diameter once it is at the desiredlocation. Mechanical intervention with stents has reduced the rate ofrestenosis as compared to balloon angioplasty. Yet, restenosis remains asignificant problem. When restenosis does occur in the stented segment,its treatment can be challenging, as clinical options are more limitedthan for those lesions that were treated solely with a balloon.

Stents are used not only for mechanical intervention but also asvehicles for providing biological therapy. Biological therapy usesmedicated stents to locally administer a therapeutic substance.Effective concentrations at the treated site require systemic drugadministration which often produces adverse or even toxic side effects.Local delivery is a preferred treatment method because it administerssmaller total medication levels than systemic methods, but concentratesthe drug at a specific site. Local delivery thus produces fewer sideeffects and achieves better results.

A medicated stent may be fabricated by coating the surface of either ametallic or polymeric scaffolding with a polymeric carrier that includesan active or bioactive agent or drug. Polymeric scaffolding may alsoserve as a carrier of an active agent or drug.

The stent must be able to satisfy a number of mechanical requirements.The stent must be capable of withstanding the structural loads, namelyradial compressive forces, imposed on the stent as it supports the wallsof a vessel. Therefore, a stent must possess adequate radial strength.Radial strength describes the external pressure that a stent is able towithstand without incurring clinically significant damage. Additionally,a stent should be sufficiently rigid to adequately maintain its size andshape throughout its service life despite the various forces that maycome to bear on it, including the cyclic loading induced by the beatingheart. For example, a radially directed force may tend to cause a stentto recoil inward. Furthermore, the stent should possess sufficienttoughness or resistance to fracture from stress arising from crimping,expansion, and cyclic loading.

Some treatments with implantable medical devices require the presence ofthe device only for a limited period of time. Once treatment iscomplete, which may include structural tissue support and/or drugdelivery, it may be desirable for the stent to be removed or disappearfrom the treatment location. One way of having a device disappear may beby fabricating the device in whole or in part from materials that erodeor disintegrate through exposure to conditions within the body. Thus,erodible portions of the device can disappear or substantially disappearfrom the implant region after the treatment regimen is completed. Afterthe process of disintegration has been completed, no portion of thedevice, or an erodible portion of the device will remain. In someembodiments, very negligible traces or residue may be left behind.Stents fabricated from biodegradable, bioabsorbable, and/or bioerodablematerials such as bioabsorbable polymers can be designed to completelyerode only after the clinical need for them has ended.

However, there are potential shortcomings in the use of polymers as amaterial for stents. For example, the mechanical properties and otherproperties are susceptible to degradation during processing.

SUMMARY OF THE INVENTION

Various embodiments of the present invention include a method offabricating a stent, comprising: forming a tube comprising abioabsorbable polymer, wherein the bioabsorbable polymer crosslinks whenexposed to radiation; forming a stent from the tube; and exposing thestent to radiation sufficient to crosslink the bioabsorbable polymer.

Additional embodiments of the present invention include a method offabricating a stent, comprising: forming a tube comprising abioabsorbable polymer, wherein the bioabsorbable polymer crosslinks whenexposed to radiation; forming a stent from the tube; and exposing thestent to radiation sufficient to crosslink the bioabsorbable polymer,wherein the bioabsorbable polymer is a copolymer of degradablefunctional groups and highly reactive functional groups, wherein thedegradable functional groups are derived from monomers that formbiodegradable polymers, wherein the highly reactive functional groupsare derived from a lactone functionalized with an alkene or an alkynegroup, wherein the radiation induces crosslinking between the degradablefunctional groups and the highly reactive functional groups.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a stent.

DETAILED DESCRIPTION OF THE INVENTION

Various embodiments of the present invention relate to implantablemedical devices, such as a stents, made from crosslinked bioabsorbablepolymers. The embodiments further relate to methods of making thedevices that include crosslinking the bioabsorbable polymer throughradiation exposure. The embodiments are generally applicable to anytubular polymeric implantable medical device. In particular, the methodscan be applied to tubular implantable medical devices such asself-expandable stents, balloon-expandable stents, and stent-grafts.

A stent may include a pattern or network of interconnecting structuralelements or struts. FIG. 1 depicts a view of a stent 100. In someembodiments, a stent may include a body, backbone, or scaffolding havinga pattern or network of interconnecting struts or structural elements105. Stent 100 may be formed from a tube (not shown). The structuralpattern of the device can be of virtually any design. The embodimentsdisclosed herein are not limited to stents or to the stent patternillustrated in FIG. 1. The embodiments are easily applicable to otherpatterns and other devices. The variations in the structure of patternsare virtually unlimited. A stent such as stent 100 may be fabricatedfrom a tube by forming a pattern with a technique such as lasermachining or chemical etching.

A stent such as stent 100 may be fabricated from a polymeric tube or asheet by rolling and bonding the sheet to form the tube. A tube or sheetfor making a stent is conventionally formed by extrusion or injectionmolding. A stent pattern, such as the one pictured in FIG. 1, can beformed in a tube or sheet with a technique such as laser cutting orchemical etching. The stent can then be crimped on to a balloon orcatheter for delivery into a bodily lumen.

An implantable medical device can be made partially or completely from abiodegradable, bioabsorbable, or biostable polymer. A polymer for use infabricating an implantable medical device can be biostable,bioabsorbable, biodegradable or bioerodable. Biostable refers topolymers that are not biodegradable. The terms biodegradable,bioabsorbable, and bioerodable are used interchangeably and refer topolymers that are capable of being completely degraded and/or erodedwhen exposed to bodily fluids such as blood and can be graduallyresorbed, absorbed, and/or eliminated by the body. The processes ofbreaking down and absorption of the polymer can be caused by, forexample, hydrolysis and metabolic processes.

A stent made from a biodegradable polymer is intended to remain in thebody for a duration of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished. Afterthe process of degradation, erosion, absorption, and/or resorption hasbeen completed, no portion of the biodegradable stent, or abiodegradable portion of the stent will remain. In some embodiments,very negligible traces or residue may be left behind.

The duration of a treatment period depends on the bodily disorder thatis being treated. In treatments of coronary heart disease involving useof stents in diseased vessels, the duration can be in a range from abouta month to a few years. However, the duration is typically up to aboutsix months, twelve months, eighteen months, or two years. In somesituations, the treatment period can extend beyond two years. The stentis expected to be completely degraded away from the vessel at the end ofthe treatment period.

As indicated above, a stent has certain mechanical requirements such ashigh radial strength, high modulus, high fracture toughness, and highfatigue resistance. A stent that meets such requirements greatlyfacilitates the delivery, deployment, and treatment of a diseasedvessel. A polymeric stent with inadequate mechanical properties canresult in mechanical failure or recoil inward after implantation into avessel.

With respect to radial strength, the strength to weight ratio ofpolymers is usually smaller than that of metals. To compensate for this,a polymeric stent can require significantly thicker struts than ametallic stent, which can result in an undesirably large profile.

Additionally, polymers that are sufficiently rigid to support a lumen atconditions within the human body may also have low fracture toughnesssince they may exhibit a brittle fracture mechanism. For example, theseinclude polymers that have a glass transition temperature (Tg) abovehuman body temperature (Tbody), which is approximately 37° C. Suchpolymers may exhibit little or no plastic deformation prior to failure.It is important for a stent to be resistant to fracture throughout therange of use of a stent, i.e., crimping, delivery, deployment, andduring a desired treatment period. PLLA is but one example of the classof semicrystalline polymers for which the above description is true. TheTg of PLLA has been reported to vary between approximately 55 and 65° C.Medical Plastics and Biomaterials Magazine, March 1998.

Certain embodiments of the present invention include a stent having astent body made from a crosslinked bioabsorbable polymer. Embodimentscan include generally a device body made from the crosslinkedbioabsorbable polymer. Additionally, embodiments of the presentinvention further include making the stent body, or more generally, thedevice body. Embodiments of the method can include forming a construct,such as a tube that includes a bioabsorbable polymer. The bioabsorbablepolymer may be uncrosslinked. Alternatively, the bioabsorbable polymermay already have some crosslinks. The bioabsorbable polymer of theconstruct becomes crosslinked when exposed to radiation. Thebioabsorbable polymer may be exposed to radiation sufficient tocrosslink the bioabsorbable polymer.

In exemplary embodiments, the bioabsorbable polymer can be a homopolymeror a copolymer. The bioabsorbable polymer can also be a polymer blend oftwo or more different types of polymer, either a miscible polymer blendor an immiscible polymer blend. The crosslinking in a polymer blend mayresult linking or bonding between polymers of different types.Alternatively, the crosslinking can be selective, in that only one typeof polymer is crosslinked or only certain types of polymers arecrosslinked in the blend.

Exemplary bioabsorbable polymers include poly(L-lactide) (PLLA),poly(D-lactide) (PDLA), polyglycolide (PGA), polymandelide (PM),polycaprolactone (PCL), poly(trimethylene carbonate) (PTMC),polydioxanone (PDO), poly(4-hydroxy butyrate) (PHB), and poly(butylenesuccinate) (PBS), poly(DL-lactide) (PDLLA), andpoly(L-lactide-co-glycolide) (PLGA). Polymers that are preferred for astent body are those that have thermal stability in the range at orclose to Tbody, since such polymer may be rigid and maintain a highmodulus and compressive strength at Tbody so that the stent can supporta lumen. Such polymers have a Tg above human body temperature,preferably at least 10, 20, or 30° C. greater than human bodytemperature. PLLA and PLGA are examples of such polymers.

In additional embodiments, the bioabsorbable polymer can also be a blendof PLLA and PDLA to create a stereocomplex, which is expected to furtherenhance the thermal and mechanical stability of the polymer. Exemplaryblends can have a ratio of PDLA to PLLA of between 0 and 1, although theblends can have a ratio greater than one.

The degree of crosslinking may be characterized by crosslink orcrosslinking density. The crosslink density can be expressed as theaverage molecular weight (number average or weight average) betweencrosslink sites (Mc). Alternatively, the crosslink density can beexpressed as the mole fraction of monomer units which are crosslinkpoints (Xc). An Introduction to Plastics, Hans-Georg Elias 2^(nd) ed.Wiley (2003). Crosslink density, the molecular weigh between crosslinks(Mc), can be determined by known methods such as dynamical mechanicalanalysis (DMA).

The crosslink density can further be described as gel fractions. The gelfraction is calculated by the amount of insoluble material in solvent,when the crosslinked polymer is mixed with a solvent for theuncrosslinked polymer, with the following equation:

Gel fraction (%)=(Wg/W ₀)100

where W₀ and Wg are the dried weights of the initial polymer and itsremaining weight (the gel component which corresponds to the crosslinkedcomponent) after dissolution in a solvent at room temperature. When acrosslinked polymer is mixed into a solvent, a portion which issufficiently crosslinked swells rather than dissolves in the solvent. Aportion which is not crosslinked or not sufficiently crosslinkeddissolves in the solvent. Therefore, the crosslinked portion may beseparated from the remainder of the polymer so that a gel fractionprovides a measure of the degree of crosslinking.

In embodiments of the present invention, the crosslink density of thecrosslinked bioabsorbable polymer of the stent body can be determinedfrom a number of techniques including equilibrium swelling (also knownas degree of swelling), NMR spectroscopy, dynamic mechanical analysis,and gel fraction. Gel fraction is typically used and should be at least1%. Sometimes, crosslink density and gel fraction are usedinterchangeably since gel fraction is related to crosslink density. Forthe purposes of this disclosure, these terms will be usedinterchangeably. The crosslink density of the bioabsorbable polymer ofthe tube after radiation exposure as determined by its gel fraction isincreased by at least 1%. More specifically, the gel fraction is or isincreased to between 1-5%, 5-20%, 20-50%, 50-70%, or greater than 70%.

Additionally, embodiments of the present invention further includemaking the stent body, or more generally, the device body. Embodimentsof the method can include forming a construct, such as a tube thatincludes a bioabsorbable polymer. The bioabsorbable polymer of theconstruct is crosslinked when exposed to radiation.

As described in more detail below, the crosslinking can be caused orinduced partially or completely by the presence of a crosslinking agentmixed or dispersed within the polymer. Alternatively or additionally,the crosslinking can be caused or induced by chemical reaction andbonding between reactive moieties present on different polymer chains ofthe bioabsorbable polymer. In this alternative, the crosslinking isbonding between functional groups of the polymer without being linked orbonded by a crosslinking agent that is distinct or separate from thepolymer prior to the crosslinking.

The crosslinked bioabsorbable polymer can result in a stent body withhigh strength sufficient to support a bodily lumen for a desired timeperiod, with high fracture toughness, and with acceptable recoil afterdeployment (less than 10% of the deployed diameter). The crosslinkedbioabsorbable polymer may have a relatively low crystallinity, forexample, between 10-25%. A distinct advantage of a stent body made froma crosslinked bioabsorbable polymer is that the polymer can have arelatively low crystallinity (e.g., 10-25%) while providing sufficientradial strength to support blood a vessel (e.g., less than 10% recoilfor at least 1-3 months) and yet have relatively high fracture toughness(e.g., few or no cracked structural elements upon deployment). Thecrystallinity can be greater than 25%, however, it is important that thedegree of crosslinking or crosslink density is not high enough to causebrittle behavior that results in unacceptable fracture or failure duringuse of a stent.

In some embodiments, the tube is exposed to radiation which causes thecrosslinking of the polymer. A stent body is made from the exposed andcrosslinked tube. In other embodiments, a stent body is fabricated fromthe tube prior to radiation exposure and the stent body is exposed tothe radiation to crosslink the bioabsorbable polymer.

The fabrication of a stent from a tube may include additional processingsteps. In some embodiments, the polymer tube can be radially deformed orexpanded, axially deformed, or both radially and axially deformed. Thestent body can be formed from the radially deformed, axially deformed,or radially and axially deformed tube. The deformation tends to increasethe strength and toughness of the polymer. In particular, the radialdeformation tends to increase the radial strength of the tube. Theincrease in radial strength is believed to be due to the circumferentialpolymer chain orientation and an increase in crystallinity, both inducedby the deformation. Both radial and axial deformation provide biaxialorientation of the polymer chains.

The radial deformation can be accomplished by a blow molding process. Insuch a process, the polymer tube is disposed within a cylindrical moldwith a diameter greater than the polymer tube. The polymer tube isheated, preferably so that its temperature is above its Tg. The pressureinside of the tube is increased to cause radial expansion of the tube sothe outside surface of the tube conforms to the inside surface of themold. The polymer tube can be axially deformed by a tensile force alongthe tube axis before, during, and/or after the radial deformation. Thepolymer tube is than cooled below Tg and further processing steps canthen be performed, such as laser machining of the tube to form a stentpattern.

The degree of radial expansion or deformation may be quantified bypercent radial expansion:

$\left\lbrack {\frac{{Outside}\mspace{14mu} {Diameter}\mspace{14mu} {of}\mspace{14mu} {Deformed}\mspace{14mu} {Tube}}{{Original}\mspace{14mu} {outside}\mspace{14mu} {Diameter}\mspace{14mu} {of}\mspace{14mu} {Tube}} - 1} \right\rbrack \times 100\%$

In some embodiments, percent radial expansion can be 200-500%. In anexemplary embodiment, the percent radial expansion is about 300%.Similarly, the degree of axial deformation may be quantified by thepercent axial elongation:

Preliminary data suggests that 200 percent radial combined with 200%axial provides the best results.

$\left\lbrack {\frac{{Length}\mspace{14mu} {of}\mspace{14mu} {Deformed}\mspace{14mu} {Tube}}{{Original}\mspace{14mu} {Length}\mspace{14mu} {of}\mspace{14mu} {Tube}} - 1} \right\rbrack \times 100\%$

In some embodiments, the tube can be elongated before during or afterthe radial expansion. The percent axial elongation can be 30-100%.

In some embodiments, the tube is radially expanded prior to the exposingstep and the stent body is formed from the expanded and exposed tube.Exposing the construct in the expanded state may be preferable since thecrosslinking may tend to reduce or inhibit recoil and improve otherphysical properties, such as toughness.

In other embodiments including the radial expansion step, the tube canbe crosslinked prior to the radial expansion step. In this embodiment, astent body fabricated from an expanded tube may have a greater tendencyto recoil toward the diameter of the tube prior to radial expansion.Recoil in this manner may be desirable for a treatment in which recoilfrom a deployed diameter is acceptable or desirable.

In another embodiment, there is no radial deformation step. In thisembodiment, the polymer tube is exposed to radiation, followed byformation of a stent pattern. Alternatively, a stent pattern is formedin the tube, followed by radiation exposure. In either case, the stentmay then be crimped. The crimped stent may then have a tendency toself-expand and can be deployed at an implant site throughself-expansion rather than balloon expansion.

As indicated above, the stent body can be formed using laser machiningto form a stent pattern in the tube. For example, a femtosecond lasercan be used. Laser machining removes material of the tube to form thestent pattern. However, material not removed that is near a machinedsurface can be modified by energy from the laser. The modification isgenerally undesirable since mechanical properties are adverselyeffected. The modified region is referred to as a heat affected zone.

Another processing step can include forming a therapeutic coating layerover all or a portion of the stent body surface. The coating can includea therapeutic substance dispersed in a polymer.

A further processing step includes mounting the stent on a deliverydevice, for example, over a balloon on a catheter. The stent can bemounted by reducing the diameter of the stent with a crimping process sothat the stent is secured to the balloon at a reduced diameter. In someembodiments, the stent body can be crosslinked at the crimped diameter.However, as mentioned above, the stent may have a strong tendency torecoil toward the crimped diameter. Such recoil may be undesirable inmany treatment situations. However, recoil in this manner may bedesirable for a treatment in which recoil from a deployed diameter isacceptable or desirable.

A further embodiment may be to treat the tube or stent with radiationafter expansion, then crimp, and finally treat again with radiation forsterilization. The second sterilization dose may be low enough so thatthe crosslink density is not increased dramatically, so that recoil isnot impaired.

In another embodiment, the tube or stent can be treated with radiationafter expansion, then crimp, and sterilize with EtO.

In another embodiment, the tube or stent may be treated with radiationafter extrusion, omit expansion, crimp, and finally sterilize with EtO.

A medical device, such as a stent, typically undergoes sterilization toreduce the bioburden of the stent to an acceptable sterility assurancelevel (SAL). Bioburden refers generally to the number of microorganismswith which an object is contaminated. SAL is a measure of the degree ofsterilization and refers to the probability of a viable microorganismbeing present on a product unit after sterilization. There are numerousmethods of sterilizing medical devices such as stents, the most commonbeing ethylene oxide treatment and treatment with ionization radiationsuch as electron beam and gamma radiation. Generally, it is desirablefor the sterilization procedure to have little or no adverse affects onthe material properties of the stent. Stents are typically sterilized ina crimped state after packaging.

Ethylene oxide (“EtO”) sterilization is performed by exposing the deviceto gaseous ethylene oxide mixtures at elevated temperatures, at highrelieve humidy for a period of time to obtain a desired bioburden level.The elevated temperatures (e.g., a temperature between 30° C. and (Tg−5° C.)) speeds up the sterilization of the device and the dissipationof the EtO from the device. Exposure to the EtO gas mixture can resultin degradation of the mechanical properties of a polymer or distortionof the fabricated shape if the conditions are not chosen properly. Thetemperature, relative humidity, and time need to be chosen carefully tobe compatible with the polymer. If any of the conditions are too high,the polymer can lose its intended shape, form, or function as the EtO,temperature, relative humidity, and time can affect these properties. Anacceptable range may be 30° C. to 50° C., 30 to 100% relative humidity,with the lowest possible temperature and humidity preferred for PLLA. Aradiation crosslinked stent would be more resistant to physicalproperty, form, function and shape degradation after exposure to EtOsterilization conditions. Radiation sterilization is well known to thoseof ordinary skill the art. Medical devices composed in whole or in partof polymers can be sterilized by various kinds of radiation, including,but not limited to, electron beam (e-beam), gamma ray, ultraviolet,infra-red, ion beam, x-ray, and laser sterilization. A sterilizationdose can be determined by selecting a dose that provides a required SAL.A sample can be exposed to the required dose in one or multiple passes.

However, it is known that radiation can alter the properties of thepolymers being treated by the radiation. Such radiation can cause adrastic drop in molecular weight of the polymer due to chains scissionand formation of free radicals. This can lead to a more brittle materialprone to cracking during deployment (expansion). Sterilization occursafter crimping.

The present invention reduces or prevents the degradation or drasticdrop in molecular weight of the polymer of a device caused by processingconditions which cause such degradation. In particular, the reduction inmolecular weight and properties caused by sterilization are reduced orprevented by the crosslinking.

Additionally, the fracture toughness of the bioabsorbable polymer can beenhanced by the crosslinking since the crosslinking reduces the degreeof crystallinity. Quynh, Tran et al., European Polymer Journal 431779-1785 (2007). Specifically, fracture toughness is enhanced if thecrosslinking is sufficiently high, but can be reduced if thecrosslinking is to high. Thus, if the radiation dose is not sufficientlyhigh to induced the sufficient degree of crosslinking, the fracturetoughness will not be enhanced. Also, if the radiation dose is too high,a degree of crosslinking can be induced that results in brittlebehavior. Furthermore, the degree of crystallinity of a stent producedby the several possible crosslinking processes described above is lowerthan one that is not crosslinked. Thus, it is expected that crackingduring crimping and deployment of the stent will be reduced oreliminated. Additionally, the crosslinking is expected to reduce thephysical aging of the polymer since the degrees of freedom in themobility of the polymer chains in reduced in the amorphous regions.

Furthermore, the crosslinking is also expected to enhance the thermaland mechanical stability of the polymer. The crosslinking tends toincrease the Tg and increase the modulus. Thus, the recoil at Tbody isreduced.

Additionally, the stent may be sterilized using EtO, which may be moredesirable than radiation since EtO may better preserve the properties ofthe finished good, i.e., the stent and catheter assembly. It has beenobserved in practice that EtO sterilization cycles tend to causebioabsorbable stents to fracture upon deployment if the stents are notcrosslinked.

As used herein, crosslinks refer generally to chemical covalent bondsthat link one polymer chain to another. A crosslinked polymer includescrosslinks throughout a polymer material sample. When polymer chains arelinked together by crosslinks, they lose some of their ability to moveas individual polymer chains, thus stabilizing the polymer.

Crosslinks can be formed by chemical reactions that are initiated byheat, pressure, crosslinking agent, or radiation. The radiation caninclude, but is not limited to, electron beam, gamma, or UV light. Thecrosslinking induced by radiation can be caused by or facilitated by acrosslinking agent. A crosslinking agent is a substance or compound thatpromotes or regulates intermolecular covalent bonding between polymerchains, linking them together to create a more rigid structure. Thecrosslinking agent is a compound that is separate and distinct from thepolymer chains prior to the crosslinking between which it promotes orregulates bonding. In its role in promoting or regulating covalentboding, the crosslinking agent becomes covalently bonded to the polymerchains. Therefore, the crosslinking agent can become incorporated intothe crosslinked polymer.

Radiation crosslinking of polylactic acids with crosslinking agents hasbeen described, for example, in Mitomo, Hiroshi et al., Polymer 464695-4703 (2005); Quynh, Tran et al., European Polymer Journal 431779-1785 (2007); and Quynh, J. of Applied Polymer Science, 110,2358-2365 (2008), which are all incorporated by reference herein. Theradiation dose, type, and mole or weight percent of a crosslinking agentcan also influence the crosslink density. The radiation dose is directlyproportional to the crosslink density.

In certain embodiments of the invention, the polymer construct that isto be irradiated, such as a tube, can include a crosslinking agent. Thecrosslinking agent can be mixed or dispersed within the bioabsorbablepolymer of the tube. When the tube is exposed to radiation, thecrosslinking agent induces crosslinking of the bioabsorbable polymer.

As already indicated, the degree of crosslinking depends on the weightpercent of the crosslinking agent and the radiation dose. The tube mayinclude an amount of crosslinking agent sufficient to provide a desiredcrosslink density or gel fraction. In exemplary embodiments, the tubeincludes less than 1 wt %, 1-3 wt %, 3-5 wt %, or greater than 5 wt %crosslinking agent. The remaining material of the tube can be thebioabsorbable polymer or consist essentially of the bioabsorbablepolymer. The tube can also include a filler material mixed with thebioabsorbable polymer and crosslinking agent.

A limiting factor on the radiation dose and amount of crosslinkingagents is that the crosslink density should not be so high that thebioabsorbable exhibits brittle fracture behavior during use of thestent, e.g., during crimping and deployment. Additionally, theconcentration of crosslinking agent in the polymer can become so highthat the performance and properties of bioabsorbable polymer arecompromised. Thus, in general, the weight percent of crosslinking agentis preferably below 5 wt %. However, there may be polymer, crosslinkingagent, and radiation dose combinations in which concentrations above 5wt % that would be favorable.

The amount of crosslinking agent and radiation dose can be varied toobtain a desired crosslink density and gel fraction. Also, the amount ofcrosslinking agent and radiation dose can be varied to obtain desiremechanical properties such as high radial strength and high fracturetoughness. The radiation dose can be 10-100 kGy, 30-40 kGy, or morenarrowly 25-30 kGy. Exemplary crosslinking agents include triallylisocyanurate (TAIC), trimethally isocyanurate (TMAIC), andtrimethylolpropane triacrylate (TMPTA), however, other crosslinkingagents may be used.

A crosslinking agent can be mixed or dispersed into the bioabsorbablepolymer of a tube using melt processing. For example, the crosslinkingagent can be fed into an extruder that in the manufacture of the tube.Alternatively, the crosslinking agent can be mixed with a polymer meltin batch and fed into a extruder or injection molder to make the tube.For example, the structure of TAIC is:

Although other crosslinking mechanisms are possible, the mechanism ofthe crosslinking can include chain scission induced by radiation betweena C—C bond in the bioabsorbable polymer, such as PLLA, which contains acarbonyl and methyl group, but can occur elsewhere as well. During chainscission, a free radical is formed. The site of the free radical canthen react with the C's of the double bond of the TAIC to form asaturated C-C crosslink.

The most effective crosslinking agent based on the magnitude of the gelfraction can be determined by measuring gel formation at variousradiation dose levels. It has been found that TAIC is the most effectivecrosslinking agent in PLLA since it exhibits the highest gel fractionfor radiation doses between 20 and 100 kGy at 3 wt % for eachcrosslinking agent. Mitomo, Hiroshi et al., Polymer 46 4695-4703 (2005).Likewise, the most effective concentration for a given crosslinkingagent based on the magnitude of the gel fraction can be determined bymeasuring gel formation at various concentration levels and radiationdoses. It has been found that a concentration of 3 wt % TAIC is the mosteffective in PLLA since it provides the highest gel fraction based on acomparison of concentrations between 0.5 wt % and 5 wt %. Mitomo,Hiroshi et al., Polymer 46 4695-4703 (2005)

Additionally, the concentration of crosslinking agent and radiation dosecan be selected based on mechanical properties such as tensile strengthand elongation at break (a measure of the fracture toughness of thepolymer). It has been shown for irradiated PLLA containing TAIC samplesthat the tensile strength was the highest for 30 kGy dose compared toother doses between 0 and 50 kGy. Id. Additionally, it has been foundthat for a PLLA/PDLA blend with a 1:1 ratio of the polymers containingTAIC, the blend with the concentration of 3 wt % TAIC showed the bestmechanical properties over a range of doses between 30 and 100 kGy.

A given radiation dose results in a corresponding degree ofcrosslinking. It may be desirable to minimize the residual unreactedcrosslinking agent present in the finished stent. Therefore, in someembodiments, the amount of crosslinking agent in the tube prior tocrosslinking can be selected so that all or substantially all of thecrosslinking agents reacts to form the crosslinks.

Alternatively, the amount of crosslinking agent can be selected so thatunreacted crosslinking agent remains in the bioabsorbable polymer afterradiation exposure. When the stent is exposed to radiation in asterilization step, additional crosslinking will occur which can furtherreduce or completely deplete the remaining crosslinking agent. This canbe advantageous since the crosslinking reaction inhibits undesirablechains scission caused by radiation. Quynh, J. of Applied PolymerScience, 110, 2358-2365 (2008).

In further embodiments of the invention, the polymer construct that isto be irradiated, such as a tube, can be composed in whole or in part ofa polymer that is crosslinkable due to formation of links or bondsbetween different moieties or functional groups of the polymer, referredto as a self-crosslinkable polymer, when exposed to radiation. In suchembodiments, the polymer crosslinks form in the absence of acrosslinking agent. In some embodiments, the polymer is free ofcrosslinking agents that are not chemically bound to the polymer chains.The crosslinking can be due entirely to crosslinking between themoieties or functional groups of the polymer. In other embodiments, thepolymer can additionally include a crosslinking agent so thatcrosslinking is due to the crosslinking agent and the reaction betweenthe moieties or functional groups of the polymer without the aid of thecrosslinking agent.

In some embodiments, the self-crosslinkable polymer can be a copolymerthat includes reactive functional groups, for example, alkenes oralkynes, and functional groups that form biodegradable polymers whenpolymerized or copolymerized. The latter functional groups (referred toas degradable functional groups) are derived from monomers that include,but are not limited to L-lactic acid, glycolic acid, caprolactone,dioxanone, D-lactic acid, mandelic acid, trimethylene carbonate,4-hydroxy butyrate, and butylene succinate. “Reactive” refers to uponexposure of the polymer to radiation, crosslinking is induced at thereactive functional groups.

The self-crosslinkable polymer can be formed through copolymerization ofcompounds that have the reactive functional groups and a monomer, suchas lactic acid, to form a biodegradable, crosslinkable polymer. Theself-crosslinkable copolymer can be a random or alternating copolymer.

Although other crosslinking mechanisms are possible, the mechanism ofthe crosslinking of the self-crosslinkable polymer can include chainscission induced by radiation between a C—C bond in a degradablefunctional groups, such as lactic acid, which contains carbonyl and themethyl group, but can occur elsewhere as well. During chain scission, afree radical is formed. The site of the free radical can then react witha reactive site of the highly reactive functional groups to form a C-Ccrosslink.

Although other types of reactive functional groups are possible, in someembodiments, the reactive functional groups include alkenes or alkynes.The double or triple bonds of the alkene or alkyne, respectively, act asthe reactive sites that form crosslinks with the degradable functionalgroups. In particular, the site of the free radical can then react withthe alkene to form a saturated C-C crosslink or react with the alkyne toform a C-C crosslink with a double bond.

In such embodiments, the bioabsorbable polymer is a copolymer formedthrough copolymerization of one or more monomers, such as lactic acid,with an alkene or alkyne. A general form of such a self-crosslinkablecopolymer is A×By, where A is a moiety such as lactic acid, and B is analkene or alkyne that can copolymerize with A, and where x is the mole %of A and y is the mole % of B in the copolymer. Exemplary compositionscan include x as 90-96 wt % and y as 4-10 wt %, although x and y can beoutside these ranges.

For example, a self-crosslinkable polymer formed from L-lactide anda-allyl-δ-valerolactone is:

Additionally, a self-crosslinkable polymer formed from L-lactide anda-alkyne-δ-valerolactone is:

In exemplary embodiments, the reactive functional groups can includelactones with alkene or alkyne groups. Exemplary alkene monomers thatcan be copolymerized with a monomers such as lactic acid to form aself-crosslinkable polymer include, but are not limited to,a-allyl-δ-valerolactone (AVL) or a-allyl-ε-caprolactone. Exemplaryalkyne monomers that can be copolymerized with a monomer such as lacticacid to form a self-crosslinkable polymer including, but are not limitedto, a-alkyne-δ-valerolactone or a-alkyne-ε-caprolactone. δ-valerolactonecan be converted to a-alkene-δ-valerolactone ora-alkyne-δ-valerolactone. Similarly, δ-caprolactone can be converted toa-alkene-δ-valerolactone or a-alkyne-δ-valerolactone. The copolymersnamed above are all random copolymers.

Schemes for functionalizing lactones with allyl and alkyne groups havebeen disclosed. Parrish, Bryan et al., J. of Polymer Science: Part A:Polymer Chemistry, Vol. 40, 1983-1990 (2002) and Parrish, Bryan et al.,J. Am. Chem. Soc., 127, 7404-7410 (2005), which are incorporated byreference herein. Lactones can be functionalized with allyl groups a tothe carbonyl as intermediates. This α methylene group is susceptible tofunctionalization under anionic conditions because of the enhancedacidity of its protons relative to the other protons on the ring. Forexample, δ-valerolactone can be functionalized with an allyl to forma-allyl-δ-valerolactone according to the following scheme (Id.):

The δ-valerolactone is quenched with allyl bromide/HMPA. Similarly,δ-valerolactone can be functionalized with an alkyne to forma-alkyne-δ-valerolactone according to the following scheme (Id.):

Both the alkene and alkyne have high thermal stability, with the alkenebeing more thermally stable than the alkyne. The high thermal stabilityallows processing with reduced or no degradation through conventionalmelt processing, such as extrusion.

A copolymer of the functionalized lactone and a degradable functionalgroup, such as L-lactic acid, may be synthesized according to thefollowing exemplary reaction between L-lactic acid anda-allyl-δ-valerolactone to formpoly(L-lactide-co-a-allyl-δ-valerolactone) (poly(LLA-co-AVL):

Additionally, the self-crosslinkable polymer can be a copolymer ofα,α-diallyl-δ-valerolactone (DAVL) and a degradable functional group,such as L-lactic acid. DAVL can be synthesized as follows:

L-lactic acid and DAVL are copolymerized to form poly(LLA-co-DAVL) asfollows:

In other embodiments, a self-crosslinkable polymer can be made through ascheme including a transesterification reaction between a degradablepolyester, such as PLLA, and a diol like PEG or a triol such as1,1,1-tris(hydroxymethyl)ethane. Then a chain extension is conductedwith the degradable polymer and an alkyne such as glycidyl propargylether or alkyne valerolactone. For example, a transesterification ofPLLA and a diol like 1,1,1-tris(hydroxymethyl)ethane can be performed asfollows:

The chain extension with the PLLA and an alkyne such as glycidylpropargyl ether or alkyne valerolactone is as follows:

In the structure above, the pendant groups containing the dots representalkyne or alkene functional groups. The alkyne monomers, glycidylpropargyl ether and alkyne valerolactone, respectively, are shown below:

The body, scaffolding, or substrate of a stent may be primarilyresponsible for providing mechanical support to walls of a bodily lumenonce the stent is deployed therein. A stent body, scaffolding, orsubstrate, for example, as pictured in FIG. 1, can refer to a stentstructure with an outer surface to which no coating or layer of materialdifferent from that of which the structure is manufactured. If the bodyis manufactured by a coating process, the stent body can refer to astate prior to application of additional coating layers of differentmaterial. “Outer surface” refers to any surface however spatiallyoriented that is in contact with bodily tissue or fluids. A stent body,scaffolding, or substrate can refer to a stent structure formed by lasercutting a pattern into a tube or a sheet that has been rolled into acylindrical shape.

In some embodiments, the stent body, scaffolding, struts, or structuralelements of the present invention may be nonporous or substantiallynonporous. Substantially nonporous refers to a porosity of less than 0.1percent. Alternatively, the stent body, scaffolding, struts, orstructural elements of the present invention may be porous.Additionally, the surface of the stent body, scaffolding, struts, orstructural elements of the present invention may have cavities oralternatively, be cavity-free.

For the purposes of the present invention, the following terms anddefinitions apply:

“Molecular weight” can refer to the molecular weight of individualsegments, blocks, or polymer chains. “Molecular weight” can also referto weight average molecular weight or number average molecular weight oftypes of segments, blocks, or polymer chains.

The number average molecular weight (Mn) is the common, mean, average ofthe molecular weights of the individual segments, blocks, or polymerchains. It is determined by measuring the molecular weight of N polymermolecules, summing the weights, and dividing by N:

${\overset{\_}{M}}_{n} = \frac{\sum_{i}{N_{i}M_{i}}}{\sum_{i}N_{i}}$

where Ni is the number of polymer molecules with molecular weight Mi.The weight average molecular weight is given by

$\mspace{79mu} {{\overset{\_}{M}\text{?}} = \frac{\sum_{i}{N_{i}M_{i}^{2}}}{\sum_{i}{N_{i}M_{i}}}}$?indicates text missing or illegible when filed

where Ni is the number of molecules of molecular weight Mi.

“Ambient temperature” can be any temperature including and between 20°C. and 30° C.

The “glass transition temperature,” Tg, is the temperature at which theamorphous domains of a polymer change from a brittle vitreous state to asolid deformable or ductile state at atmospheric pressure. In otherwords, the Tg corresponds to the temperature where the onset ofsegmental motion in the chains of the polymer occurs. When an amorphousor semicrystalline polymer is exposed to an increasing temperature, thecoefficient of expansion and the heat capacity of the polymer bothincrease as the temperature is raised, indicating increased molecularmotion. As the temperature is raised the actual molecular volume in thesample remains constant, and so a higher coefficient of expansion pointsto an increase in free volume associated with the system and thereforeincreased freedom for the molecules to move. The increasing heatcapacity corresponds to an increase in heat dissipation throughmovement. Tg of a given polymer can be dependent on the heating rate andcan be influenced by the thermal history of the polymer. Furthermore,the chemical structure of the polymer heavily influences the glasstransition by affecting mobility.

“Toughness” is the amount of energy absorbed prior to fracture, orequivalently, the amount of work required to fracture a material. Onemeasure of toughness is the area under a stress-strain curve from zerostrain to the strain at fracture. The units of toughness in this caseare in energy per unit volume of material. See, e.g., L. H. Van Vlack,“Elements of Materials Science and Engineering,” pp. 270-271,Addison-Wesley (Reading, Pa., 1989).

The underlying structure or substrate of an implantable medical device,such as a stent can be completely or at least in part made from abiodegradable polymer or combination of biodegradable polymers, abiostable polymer or combination of biostable polymers, or a combinationof biodegradable and biostable polymers. Additionally, a polymer-basedcoating for a surface of a device can be a biodegradable polymer orcombination of biodegradable polymers, a biostable polymer orcombination of biostable polymers, or a combination of biodegradable andbiostable polymers.

It is understood that after the process of degradation, erosion,absorption, and/or resorption has been completed, no part of the stentwill remain or in the case of coating applications on a biostablescaffolding, no polymer will remain on the device. In some embodiments,very negligible traces or residue may be left behind. For stents madefrom a biodegradable polymer, the stent is intended to remain in thebody for a duration of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished.

Other representative examples of polymers that may be used to fabricatean implantable medical device include, but are not limited to,poly(N-acetylglucosamine) (Chitin), Chitosan, poly(hydroxybutyrate),poly(hydroxybutyrate-co-valerate), polyorthoester, polyanhydride,polyester amide, poly(glycolic acid-co-trimethylene carbonate),co-poly(ether-esters) (e.g. PEO/PLA), polyphosphazenes, biomolecules(such as fibrin, fibrinogen, cellulose, starch, collagen and hyaluronicacid), polyurethanes, silicones, polyesters, polyolefins,polyisobutylene and ethylene-alphaolefin copolymers, acrylic polymersand copolymers other than polyacrylates, vinyl halide polymers andcopolymers (such as polyvinyl chloride), polyvinyl ethers (such aspolyvinyl methyl ether), polyvinylidene halides (such as polyvinylidenechloride), polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics(such as polystyrene), polyvinyl esters (such as polyvinyl acetate),acrylonitrile-styrene copolymers, ABS resins, polyamides (such as Nylon66 and polycaprolactam), polycarbonates, polyoxymethylenes, polyimides,polyethers, polyurethanes, rayon, rayon-triacetate, cellulose, celluloseacetate, cellulose butyrate, cellulose acetate butyrate, cellophane,cellulose nitrate, cellulose propionate, cellulose ethers, andcarboxymethyl cellulose. Another type of polymer based on poly(lacticacid) that can be used includes graft copolymers, and block Copolymers,such as AB block-copolymers (“diblock-copolymers”) or ABAblock-copolymers (“triblock-copolymers”), or mixtures thereof.

Additional representative examples of polymers that may be especiallywell suited for use in fabricating or coating an implantable medicaldevice include ethylene vinyl alcohol copolymer (commonly known by thegeneric name EVOH or by the trade name EVAL), poly(butyl methacrylate),poly(vinylidene fluoride-co-hexafluororpropene) (e.g., SOLEF 21508,available from Solvay Solexis PVDF, Thorofare, N.J.), polyvinylidenefluoride (otherwise known as KYNAR, available from ATOFINA Chemicals,Philadelphia, Pa.), ethylene-vinyl acetate copolymers, and polyethyleneglycol.

EXAMPLES

Some embodiments of the present invention are illustrated by thefollowing examples. The examples are being given by way of illustrationonly and not by way of limitation. The parameters and data are not beconstrued to unduly limit the scope of the embodiments of the invention.

Example 1

Synthesis of poly(L-lactide-co-a-allyl-δ-valerolactone)(poly(LLA-co-AVL) was performed according to the reaction schemedescribed above. All reactions were performed with 0.04 mol % Sn(Oct)2catalyst. Details of the synthesis and molecular weight data areprovided in Table 1.

TABLE 1 Poly(LLA-co-AVL) composition and molecular weight analysis. FeedIncorp. Reaction Yield [mol %] AVL^(a) Time [h] [%] AVL:LLA [mol %]Mn^(b) Mw^(b) PDI^(b) 1a 17 38  0:100 — 67k 109k 1.63 1b 18 37  0:100 —122k  221k 1.81 2 15 68  5:95 1.4 74k 125k 1.70 3 15 75 10:90 2.6 61k135k 2.20 4a 20 48 20:80 3.9 69k 115k 1.65 4b 20 73 20:80 3.7 129k  194k1.51 5 20 55 30:70 5.6 82k 124k 1.43 ^(a)Determined by 1H-NMR,^(b)Determined by GPC in CHCl3 using Polystyrene standards, Molecularweight in g/mole, PDI = polydispersity index.

Samples of poly(LLA-co-AVL) were prepared with different molepercentages of AVL. The samples were analyzed with dynamic scanningcalorimetry (DSC) to determine the Tg and the melting temperature (Tm).The results of the DSC analysis are shown in Table 2.

TABLE 2 Tg and Tm from DSC for different mole percentages of AVL inpoly(LLA-co-AVL). Mol % AVL in PLLA 0 1.4 2.6 3.9 5.6 Tg [° C.] 61 59 5852 48 Tm [° C.] 175 160 164 159 150

Example 2

Synthesis of poly(L-lactide-co-a α,α-diallyl-δ-valerolactone)(poly(LLA-co-DAVL) were performed according the to reaction schemedescribed above. All reactions were performed with 0.04 mol % Sn(Oct)2catalyst. The reactions for samples 1 and 2 were performed without aninitiator. Details of the sythesis and molecular weight data areprovided in Table 3.

TABLE 3 Poly(LLA-co-AVL) composition and molecular weight Analysis. FeedIncorp. Reaction Yield [mol %] DAVL Time [h] [%] DAVL:LLA [mol %] Mn^(b)Mw^(b) PDI^(b) 1 18 57 10:90 0.6 82k 143k 1.73 2 18 39 20:80 0.6 60k111k 1.87 3 17 45 30:70 — 49k  81k 1.64 ^(a)Determined by 1H-NMR,^(b)Determined by GPC in CHCl3 using polystyrene standards.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

1. A method of fabricating a stent, comprising: forming a tubecomprising a bioabsorbable polymer, wherein the bioabsorbable polymercrosslinks when exposed to radiation; forming a stent from the tube; andexposing the stent to radiation sufficient to crosslink thebioabsorbable polymer.
 2. The method of claim 1, further comprisingsterilizing the stent body through exposure to radiation, wherein thecrosslinking reduces or prevents molecular weight degradation of thebioabsorbable polymer due to the sterilizing radiation.
 3. The method ofclaim 1, further comprising sterilizing the stent body with ethyleneoxide at an elevated temperature above 30° C., wherein the crosslinkingreduces or prevents mechanical property degradation of the bioabsorbablepolymer due to the deleterious effect of EtO, humidity and the elevatedtemperature of the ethylene oxide sterilization.
 4. The method of claim1, further comprising radially expanding the tube prior to the formingthe stent, wherein the stent body is formed from the expanded tube. 5.The method of claim 1, wherein the stent comprises a crosslinking agentthat induces crosslinking of the bioabsorbable polymer when the tube isexposed to the radiation.
 6. The method of claim 1, wherein the stentcomprises TRIC that induces crosslinking of the polymer when the tube isexposed to the radiation.
 7. The method of claim 1, wherein theradiation exposure is 20-40 kGy.
 8. The method of claim 1, wherein thebioabsorbable polymer comprises PLLA or a PLLA and PDLA blend whichforms a stereocomplex.
 9. A method of fabricating a stent, comprising:forming a tube comprising a bioabsorbable polymer, wherein thebioabsorbable polymer crosslinks when exposed to radiation; forming astent from the tube; and exposing the stent to radiation sufficient tocrosslink the bioabsorbable polymer, wherein the bioabsorbable polymeris a copolymer of degradable functional groups and highly reactivefunctional groups, wherein the degradable functional groups are derivedfrom monomers that form biodegradable polymers, wherein the highlyreactive functional groups are derived from a lactone functionalizedwith an alkene or an alkyne group, wherein the radiation inducescrosslinking between the degradable functional groups and the highlyreactive functional groups.
 10. The stent of claim 9, wherein thedegradable functional groups are derived from L-lactic acid.
 11. Thestent of claim 9, wherein the degradable functional groups are derivedfrom monomers selected from the group consisting of glycolic acid,caprolactone, dioxanone, D-lactic acid, mandelic acid, trimethylenecarbonate, 4-hydroxy butyrate, and butylene succinate.
 12. The stent ofclaim 9, wherein the lactone is allyl-δ-caprolactone ora-diallyl-δ-caprolactone
 13. The stent of claim 9, wherein the copolymeris of the form A×By, wherein A is L-lactic acid and B is an alkene oralkyne copolymerized with A, and wherein x is the mole % of A and y isthe mole % of B in the copolymer.
 14. The stent of claim 13, wherein xis 90-96 wt % and y is 4-10 wt %.
 15. The stent of claim 13, wherein Bis selected from the group consisting of a-allyl-δ-valerolactone,a-diallyl-δ-valerolactone, a-alkyne-δ-valerolactone,a-allyl-δ-caprolactone, a-diallyl-δ-caprolactone, anda-alkyne-δ-caprolactone.